The present invention relates to dedicated radio frequency coils for use in magnetic resonance imaging (MRI) of biological subject, and more particularly tuned inductively coupled coils of this type.
The development of dedicated radio frequency (RF) coils has long been a topic of interest in the MRI field. A useful introduction to the subject can be found in the article by W. T. Sobel, "Dedicated Coils in Magnetic Resonance Imaging," Reviews of Magnetic Resonance in Medicine, Vol. 1, No. 2, pp. 181-224 (1986). In his paper, Sobel distinguishes between magnetic resonance imaging (MRI) and magnetic resonance spectroscopy (MRS). The same distinction is made in the present specification, in which references to imaging refer to two-dimensional Fourier transform imaging (which may be derived from spectral information) as opposed to spectroscopy which involves the reproduction of a nuclear magnetic resonance spectrum.
Several RF coil features and parameters directly affect its suitability for use in magnetic resonance imaging. Ideally, the coil would have a high uniform sensitivity within a particular spatial region of interest, and low sensitivity elsewhere, with a resulting high signal-to-noise (SNR) ratio. The coil would be large enough to achieve the spatial coverage desired, but small to achieve a high fill factor, to fit within the MRI system magnet and to conform comfortably to the body of a subject. The RF coil must resonate at approximately the Larmor frequency of the nuclei used to develop the MRI signal, so that neither the coil size nor geometry can create an inductance or self-capacitance which prevents tuning to the desired frequency. The coil must also couple to the MRI system amplifier stage efficiently.
One technique for coupling the RF coil to the imaging system amplifier stage is inductive coupling. In this scheme, a primary winding is positioned proximate the part of the subject which is to be imaged and a second winding, typically a single loop, is positioned adjacent the primary winding for inductively coupling with it. The secondary loop is coupled to the MRI system amplifier. Magnetic resonance signals are excited within the subject under magnetic field conditions to permit imaging and are received by the primary winding. Current flowing in the primary winding induces a voltage in the secondary winding which is amplified and processed to develop an image.
At this juncture, there is no comprehensive analysis of inductively coupled RF coils for MRI. The importance of the degree of inductive coupling between the primary and secondary coils, their relative spatial positions and respective geometries, and how these factors fit into the other aspects of RF coil design mentioned above, remain largely unanswered. The various inductively coupled RF coils analyzed in the literature appear to be special cases of the general problem.
The article by W. Froncisz et.al., "Inductive (Flux Linkage) Coupling to Local Coils in Magnetic Resonance Imaging and Spectroscopy," Journal of Magnetic Resonance 66, pp. 135-143 (1986), presents an analysis of an inductively coupled coil for use in MRI. The secondary coil is an untuned single loop and its sole purpose is to couple the primary coil to a receiver. The article concludes, among other things, that detuning can be minimized by making the secondary coil with the smallest possible inductance and coupled as tightly as possible to the primary coil.
A loop array structure was proposed by C. Leussler et.al. "Optimized RF Coils for Low Field MRI," Proc. SMRM 1989, p. 938, for applications where solenoids previously had found use. They disclose a head coil comprised of eleven turns and a body coil of eight turns, wherein each turn is a single loop LC resonator. Each resonator is tuned to the same frequency. The article presents data to show that the Q of the loop array is degraded less than the Q of the solenoid when loaded, over a frequency range of about 2.5 to 25 MHz.
Signal-to-noise ratio was considered in the article of D. J. Gilderdale et.al. in "The Performance of Mutually-Coupled Coils for Magnetic Resonance Signal Recovery," Proc. SMRM, p. 956 (1989). The authors concluded that in an inductively coupled coil system, the best SNR is obtained when the primary and secondary coils are slightly overcoupled and the lower frequency peak of the coil system frequency response is tuned to the frequency of interest.
The SNR of inductively coupled systems in which the secondary coil is significantly larger than the primary coil was investigated in the paper by S. N. Wright, "Estimation of the SNR Loss Due To Inductive Coupling Loops," Proc. SMRM, p. 955 (1989). The paper concludes that there will be less than a five percent drop in SNR of the coupled system relative to the primary alone, if the input resistance of the system is greater than ten times the primary coil resistance.
Finally, L. Darrasse et.al., "Optimization of Receiver Coil Bandwidth by Inductive Coupling," Proc. SMRM, p. 1340 (1990) discloses a strongly overcoupled inductively coupled coil having a large bandwidth for fast scanning on a low field MRI system.
Significantly, the prior art largely relies on an analysis of inductively coupled coils which is based on a lumped parameter circuit model. There is little, if any, consideration of the spatial sensitivity distribution of the primary coil, and usually there is a tacit assumption that the NMR signal is received by only the primary coil, and then inductively coupled to the secondary coil. Additionally, signal-to-noise ratio is evaluated in terms of the NMR signal voltage and the noise voltage which is output from the secondary coil; not the SNR of the image which is ultimately formed. However, the spatial sensitivity of the coil system can play a determinative role in the image SNR that is achieved.